JP2014517751A - Management of degradation profile of bioabsorbable poly (L-lactide) scaffold - Google Patents

Abstract

A method for managing a degradation profile of a biodegradable stent scaffold is disclosed. The disclosed method includes managing degradation profile characteristics including radial strength loss time and stent degradation time.

Description

  The present invention relates to a bioabsorbable polymer medical device, and more particularly to a method of treating a blood vessel with a stent.

  The present invention relates to a radially expandable endoprosthesis adapted for implantation in a body lumen. An “endoprosthesis” is a prosthesis that is placed in the body. A “lumen” refers to a cavity of a tubular organ such as a blood vessel. A stent is an example of an endoprosthesis. A stent is a generally cylindrical device that functions to keep blood vessels or other anatomical lumens such as the urethra and bile ducts open and sometimes dilated. Stents are often used in the treatment of intravascular atherosclerotic stenosis. “Stenosis” refers to a narrowed or constricted passage or opening in the body. In such treatment, the stent reinforces the blood vessels in the body and prevents restenosis after vascular angioplasty. “Restenosis” refers to a vascular or heart valve stenosis that recurs after apparently successful treatment (balloon angioplasty, stent implantation or valvuloplasty).

  Stents are typically composed of scaffolds, which contain a pattern or network of interconnected structural elements or struts (struts), and struts can be wires, tubes, or It is formed of a sheet rolled into a cylindrical shape. The scaffold was named because it is physically open and expands the walls of the aisle as needed. Typically, the stent is compressed or crimped onto the catheter so that it can be delivered to the treatment site and deployed.

  Delivery includes using a catheter to insert the stent into a narrow lumen and carry the stent to the treatment site. Deployment involves greatly expanding the diameter of the stent when it reaches the desired site. The rate of restenosis is lower in mechanical treatment with stents compared to balloon angioplasty. Still, restenosis has a major problem. If restenosis occurs at the site where the stent is implanted, it can be difficult to treat because there are limited treatment options than in the case of lesions in these tissues that are only treated with a balloon.

  Stents are used not only as mechanical treatments but also as a means of providing biological therapy. Biotherapy uses dosing stents to administer therapeutic agents locally. The therapeutic agent can also mitigate adverse biological responses to the presence of the stent. Effective concentrations at the treatment site require systemic drug administration, which often causes side effects or even toxic side effects. Local delivery is the preferred treatment because it concentrates the drug at the designated site and requires less total dosage than systemic administration. Thus, local delivery has fewer side effects and achieves better results.

  Dosage stents can be made by coating the surface of a metal or polymer scaffold with a polymeric carrier containing an active agent or bioactive agent or drug. The polymer scaffold can also function as a carrier for the active agent or drug.

  The stent must be able to meet a number of mechanical requirements. The stent must have sufficient radial strength to withstand the structural load, ie, the radial compressive force applied when the stent supports the vessel wall. The “radial strength” of a stent is defined as the pressure when it undergoes an irreversible deformation. Following the loss of radial strength, mechanical integrity gradually decreases.

  As the stent expands, the lumen must be supported accordingly during the time required for treatment, despite the various forces applied to the stent, including the cyclic load caused by the pulse. Furthermore, the stent must be sufficiently flexible with some resistance to failure.

  The treatment of coronary artery disease has undergone three changes since the 1970s. The first is a balloon angioplasty in the 1970s, followed by a metal stent in the 1990s and a third is a drug eluting stent (DES) made of metal in the 2000s. Currently, all commercially available metallic DES are made of biostable metals and remain permanently in the body after implantation, making further non-invasive screening or retreatment difficult.

  Stents are made from biostable or non-dissolvable materials such as metals and are used in both percutaneous coronary angioplasty (PCI) and peripheral blood vessels such as superficial femoral artery (SFA). Is recognized as being able to prevent recoil (rest) and restenosis early and later and has become the standard of care.

  In order to increase the therapeutic effect on the diseased blood vessels, the stent needs to exist for a limited period of time. Advances in biodissolvable stents or scaffolds can prevent permanent metal implantation in the blood vessel and allow for gradual expansion of lumens and blood vessels, and after complete absorption of the scaffold, Only healed native vascular cells can be left behind. Stents made from biodegradable, bioabsorbable, and / or biosoluble materials, such as bioabsorbable polymers, are completely dissolved only after the stent is no longer needed for treatment or after a short time Can be designed to do. As a result, fully bioabsorbable stents reduce or eliminate potential long-term complications and the risk of late thrombosis, facilitate non-invasive diagnostic MRI / CT imaging, Allows recovery and offers the possibility of plaque regression. Furthermore, the bioabsorbable stent does not permanently seize the side branch or prevent the use of future follow-up non-invasive imaging.

  Unlike durable stents, the properties of bioabsorbable stents change dramatically over time once implanted. The ability of a stent to provide adequate treatment depends not only on its initial properties, but also on its properties as a function of time, ie its degradation profile (overview). Degradation profiles, such as how long the stent can support the lumen at the deployed diameter, and the time to complete bioabsorption affect the behavior essential for proper treatment.

  In summary, a completely biosoluble scaffold has the potential to restore vascular integrity as the latest vascular recovery therapy expected to be a revolution in the fourth vascular disease therapy. While this new concept is very intriguing, so far most in vivo soluble scaffold projects developed by various companies and academia are far from actual commercialization. One of the main reasons is that many researchers in this area focus on scaffold quality control work at time zero (ie, at the time of implantation before the degradation begins in the lumen) There has been no proper approach to management.

Cross-reference of related applications, as if individual publications or patent specifications were incorporated specifically and individually by reference, and each individual publication or patent specification contained all figures, All publications and patent specifications mentioned herein are hereby incorporated by reference as if fully set forth herein.

  Various embodiments of the present invention include a method of manufacturing a bioabsorbable stent, the method comprising: providing a bioabsorbable polymer; bioabsorbable implant so that it is fully absorbed after implantation Selecting a desired degradation time range of the stent scaffold; determining a range of Mn (0) for a stent made from a bioabsorbable polymer, wherein the finished stent provides a range of degradation times; The determined Mn (0) range of the finished stent is determined from a kinetic degradation model of the bioabsorbable polymer; and making a stent scaffold from the bioabsorbable polymer, comprising: The scaffold has a Mn (0) within the determined range of Mn (0).

  Further embodiments of the present invention include a method of manufacturing a bioabsorbable stent, the method comprising: providing a bioabsorbable polymer; bioabsorbable implantable stent scaffold for provision to an implantation site Selecting a desired minimum opening time for: determining Mn at the time of loss of radial strength of a bioabsorbable stent scaffold made from a bioabsorbable polymer; equal to Mn at loss of radial strength, desired Determining the Mn (0) of a stent scaffold made from a bioabsorbable polymer that provides Mn at a minimum opening time, wherein the determined Mn (0) is the kinetics of the bioabsorbable polymer. Determined from the degradation model; and in making the stent scaffold from the bioabsorbable polymer I, stent scaffold has been determined Mn (0) or more of Mn (0).

  Further embodiments of the present invention include a method of manufacturing a bioabsorbable stent, the method comprising: providing a bioabsorbable polymer, wherein the bioabsorbable polymer is formed by a monomer polymerization reaction. Selecting a desired degradation time range of the bioabsorbable implantable stent scaffold to be completely absorbed after implantation; to provide a range of degradation time for the stent scaffold; Determining the range of monomer content within the bioabsorbable polymer, wherein the determined monomer content range is determined from a kinetic degradation model of the bioabsorbable polymer; and bioabsorbability Making a stent scaffold from a polymer, the stent scaffold being a determined range Having a monomer content.

  Another embodiment of the present invention includes a method of manufacturing a bioabsorbable stent, the method comprising: providing a bioabsorbable polymer, wherein the bioabsorbable polymer is formed by a monomer polymerization reaction. Selecting a desired minimum open time of the bioabsorbable implantable stent scaffold; Mn at the time of loss of radial strength of the bioabsorbable stent scaffold made from the bioabsorbable polymer; Determining; determining the monomer content of the bioabsorbable polymer of the finished stent that provides Mn at a desired minimum opening time equal to Mn at the loss of radial strength, the determined monomer content Is determined from a model of bioabsorbable polymer degradation characteristics versus monomer content; and bioabsorption A step of producing a stent scaffold from the polymer, a bioabsorbable polymer stent scaffold having a monomer content of less monomer content determined.

  A further embodiment of the invention includes a method of manufacturing a bioabsorbable stent, the method comprising: providing a bioabsorbable polymer scaffold made of PLLA prior to the radiation exposure step. The scaffold has a PLLA Mn of at least about 250 kDa; selecting a desired minimum opening time of the PLLA scaffold to provide to the implantation site; upon loss of radial strength during PLLA scaffold degradation Providing Mn; determining Mn (0) of PLLA scaffold that provides Mn of PLLA scaffold at a desired minimum opening time equal to Mn at loss of radial strength; and Mn of PLLA scaffold Is reduced to a value not lower than Mn (0) 3 The radiation dose ~75kGy comprising the step of exposing the PLLA scaffold performing sterilization.

  Another embodiment of the invention includes a method of making a bioabsorbable stent, the method comprising: providing a PLLA polymer scaffold, wherein the PLLA polymer tube has a Mn of at least 250 kDa; Exposing the laser-cut scaffold to a first radiation dose before crimping to reduce Mn; crimping the exposed scaffold to a reduced diameter on the delivery balloon; Exposing the fold to a second radiation dose of 20-31 kGy for sterilization to reduce Mn to Mn (0), wherein Mn (0) has a degradation time of 16-20 months and at least about 3 Time of loss of radial strength in months.

  Another embodiment of the present invention includes a method of manufacturing a bioabsorbable stent, the method comprising: providing a bioabsorbable polymer scaffold made of PLLA prior to the radiation exposure step. The scaffold PLLA Mn is at least about 250 kDa; and exposing the scaffold to radiation for sterilization, wherein the radiation reduces the scaffold Mn to 70 kDa or less and is exposed. Scaffold Mn provides a degradation time of exposed scaffold of less than 18 months and a time of at least 3 months to loss of radial strength.

  Another embodiment of the invention includes a method of manufacturing a stent, the method comprising: providing a PLLA resin having a Mn of 150-200 kDa; treating the PLLA to form a PLLA scaffold Forming a coating comprising PDLLA having a Mn of 80-100 kDa on the PLLA scaffold; and exposing the coated scaffold to radiation for sterilization, wherein the exposure to PLLA Reduce the Mn of the scaffold to 70 kDa or less.

  Another embodiment of the invention includes a method of manufacturing a stent, the method comprising: providing a bioabsorbable polymer resin; extruding the polymer resin to form a tube; polymer tube Radially expanding; creating a stent scaffold from the expanded tube; radiation sterilizing the scaffold; and at least one of a resin, an extruded tube and a radially expanded tube In order to reduce Mn in advance by hydrolysis.

  Another embodiment of the invention includes a method of manufacturing a stent, the method comprising: making a PLLA stent scaffold made of a biodegradable stent scaffold, the PLLA stent scaffold A pre-degradation step of hydrolysis of the PLLA stent scaffold prior to radiation sterilization to reduce the Mn of the PLLA stent scaffold to less than 100 kDa; The treatment provides a scaffold degradation time of less than 18 months.

FIG. 1 shows an exemplary stent scaffold. FIG. 2 is a schematic representation of the degradation behavior of a bioabsorbable scaffold overlaid with graphs of molecular weight loss, strength loss, and mass loss. FIG. 3A shows the degradation profiles of PLLA scaffolds with different monomer concentrations. FIG. 3B shows the degradation profile from in vitro testing of two lots of animal experiments. FIG. 4 is a schematic diagram of the dependence found by the inventor of the decomposition profile and its related properties with respect to Mn and the decomposition rate or decomposition rate constant. FIG. 5 shows the change in mechanical strength of the bioabsorbable scaffold when Mn is changed. FIG. 6 shows three degradation profiles, and profile 1 shows the Mn at the third month, which is equivalent to the three-month opening required for cardiac treatment. FIG. 7 shows the degradation rate constant (k) as a function of lactide content calculated from the linear regression plot of FIG. 3A. FIG. 8 shows the degradation profiles of two bioabsorbable scaffolds with different starting Mn and monomer concentrations. FIG. 9 shows the Mn versus time for the PLLA scaffold. FIG. 10 shows the Mn versus degradation time for two specifications of the PLLA scaffold. FIG. 11 shows logk (including velocity) vs. 1 / T within the relevant temperature range for PLLA samples digested in PBS buffer solution. FIG. 12 shows the normalized Mn versus time before decomposition at five stages of temperature. FIG. 13 shows the effect of the PLLA scaffold production process on monomer lactide production. FIG. 14 shows the lactide content of an extruded tube obtained from gas chromatography by a flame ionization type detection method. FIG. 15 shows the change in radial strength over the degradation time as a function of lactide content for 4 lots of extruded tubes from Example 2.

  Coronary arteries generally refer to arteries that branch off from the aorta that supply the heart muscle with oxygen-containing blood. Peripheral arteries generally refer to blood vessels other than the heart and brain. Both coronary artery disease and peripheral arterial disease cause the arteries to harden and narrow, ie, stenosis, restricting blood flow. Coronary arteries restrict blood flow to the heart, while peripheral arteries restrict blood flow to the kidneys, stomach, arms, legs, and feet. Stenosis is caused by the deposition of cholesterol and other substances on the blood vessel wall called plaque. These narrowed or narrowed portions are often called lesions. Arterial disease also includes restenosis, that is, restenosis that occurs after angioplasty treatment. There may be several mechanisms that lead to arterial restenosis, but the important one is the inflammatory response, which induces tissue growth around the site of angioplasty. Inflammatory reactions can be caused by inflation of the balloon used to open the blood vessel or, when the stent is deployed, by foreign matter on the stent itself.

  Embodiments of the present invention include treatment of various body lumens with bioabsorbable polymer stents, particularly treatment of coronary and peripheral arterial disease in the coronary arteries, and various types including the femoral superficial artery, iliac artery, and carotid artery Applicable to the treatment of peripheral blood vessels. This embodiment is further applicable to various stents such as self-expanding and balloon expandable stents. The embodiments are further applicable to a variety of stent designs including scaffold structures often formed from tubes, wire structures, and woven mesh structures.

  In embodiments of the invention, a stent can include a plurality of cylindrical rings that are coupled or coupled to a link element. When deployed in a section of a blood vessel, the cylindrical ring withstands loads in the range of expanded diameter or diameter due to periodic forces within the blood vessel and supports the vessel wall. Withstanding a load refers to supporting a load imposed by a radially inwardly directed force. Structural elements such as link elements or struts are not load bearing elements and serve to maintain the connection between the rings. For example, stents have a scaffold composed of interconnected structural elements or strut patterns, or networks.

  FIG. 1 shows an exemplary stent 100. In some embodiments, the stent includes a scaffold having a pattern or network of structural elements 105 interconnecting the body, backbone or backbone. The stent 100 may be formed from a tube (not shown). FIG. 1 shows typical features for many stent patterns including a cylindrical ring 107 connected by a link element 110. As described above, the cylindrical ring provides a radial force to support the vessel wall to withstand the load. The link elements generally function to hold the cylindrical rings together. A structure, such as stent 100, having multiple structural elements is referred to as a stent scaffold or scaffold. The scaffold may further include a coating, but it is the scaffold structure that is a load bearing structure that is responsible for supporting the lumen wall when the scaffold is expanded in the lumen.

  The structural pattern of FIG. 1 is merely exemplary and only describes the basic structure and features of the stent pattern. Stents such as stent 100 can be made by rolling and welding a polymer tube or sheet to form a tube. The tube or sheet can be formed by extrusion or injection molding. A stent pattern as shown in FIG. 1 can be formed on a tube or sheet by techniques such as laser cutting or chemical etching. The stent can then be crimped onto a balloon or catheter for delivery to the body lumen.

  The process of making a stent scaffold includes the selection of a bioabsorbable polymer material or resin. The processing steps to create the stent scaffold include resin melt processing (extrusion) to form the tube, optional tube expansion, tube laser cutting to form the scaffold, and laser-cut scaffolding. This includes optional coating on the fold, crimping the laser cut scaffold on the delivery balloon, crimping, stent and balloon packaging, and radiation sterilization steps.

  As a degradation mechanism of a biodegradable polymer, a method of chemically hydrolyzing a backbone unstable to hydrolysis is mainly used. For bulk (mostly) dissolved polymers, the polymer is chemically degraded throughout the polymer volume. As the polymer degrades, the molecular weight decreases. Following the decrease in molecular weight, mechanical properties (eg, strength) and stent properties decrease. Following the deterioration of the mechanical properties, it leads to a loss of mechanical integrity and then to dissolution or loss of mass. Mechanical integrity is demonstrated by cracking and fragmentation. Fragmental enzymatic attack and metabolism occurs, resulting in a rapid loss of polymer mass.

上式で、Ｎｉは、分子量Ｍｉのポリマー分子の個数である。重量平均分子量は次式で与えられる：

上式で、Ｎｉは、分子量Ｍｉの分子の個数である。他に指定がない限り、「分子量」は数平均分子量（Ｍｎ）を指すことにする。 The term “molecular weight” may refer to one or more definitions of molecular weight. “Molecular weight” may refer to the molecular weight of an individual segment, block, or polymer chain. “Molecular weight” may refer to the weight average molecular weight or number average molecular weight of various segments, blocks, or polymer chains.
Number average molecular weight (Mn) is the normal, average molecular weight of an individual segment, block, or polymer chain. The molecular weight is usually expressed in gram / mole called “Dalton”. This is determined by measuring the molecular weight of N polymer molecules, summing the weights and dividing by N:

In the above formula, Ni is the number of polymer molecules having a molecular weight Mi. The weight average molecular weight is given by:

In the above formula, Ni is the number of molecules having a molecular weight Mi. Unless otherwise specified, “molecular weight” refers to number average molecular weight (Mn).

  Treatment of arterial disease with the stent of the present invention will have time dependent properties when the stent is implanted to allow treatment and recovery of vascular lesions. In particular, time dependent properties include molecular weight, mechanical properties, stent properties (eg, radial strength), mechanical integrity, and mass. The treatment process can be associated with a phase of degradation characteristics schematically illustrated in FIG.

  FIG. 2 is a schematic diagram illustrating a life cycle that can be explained by a decrease in molecular weight sequence, loss of strength, and loss of mass of poly (L-lactide) scaffold after implantation in vivo. Pistner H, Bendix D, Muhling J, Reuter J. Poly (L-lactide): In vivo study of long-term degradation properties Biomaterial, 1993; 14: 291-298.

  This decomposition / absorption can be further divided into three phases. During phase I, molecular weight reduction occurs, but neither mechanical strength nor mass is affected. As the molecular weight becomes so low as to affect the mechanical properties of the scaffold, the material enters the phase II degradation properties region and the scaffold gradually loses strength. In Phase III, significant mass loss occurs after hydrolytic molecular chain breakage produces water-soluble low molecular weight species.

  Of the three phases, Phase I is particularly important for treatment with bioabsorbable scaffolds. During Phase 1, the scaffold is required to function similar to a permanent metal stent to prevent restenosis, which occurs primarily due to contractile remodeling (vasoconstriction). Ormiston JA, Serruys PW, Blood Circulation: Cardiovascular Intervention 2, 255 (2009). As described in detail herein, the inventor has determined that the duration of Phase I, ie the loss of radial strength, is two parameters (see FIG. 8): 1) Kinetic degradation characteristics (degradation rate) And 2) it was found to depend on the initial value of molecular weight (Mn (0)) at scaffold decomposition time t = 0. The reason why Mn (number average molecular weight) is adopted is that, when hydrolysis occurs in each polymer chain, it is highly related to hydrolysis. As described in detail herein, the inventors have demonstrated that the management of kinetic degradation properties leads to the lactide content in the extruded tube leading to the lactide content specification in the process. It can be achieved through management. Unless otherwise specified, lactide refers to an L-lactide monomer that is not polymerized, that is, not chemically bonded to another molecule.

  In Phase I of the degradation process, the scaffold provides an initial need for mechanical support to maintain, ie, keep open, the vascular patency at or near the deployment diameter. The patency provided by the stent allows the stented portion of the blood vessel to undergo positive remodeling at an increased deployment diameter and prevent negative remodeling. Remodeling generally refers to a structural change in the vessel wall that has an enhanced load bearing capability such that the diameter of the vessel wall where the stent is implanted remains increased without support of the stent. . The opening period is necessary to permanently acquire a positive reconfiguration.

  During Phase I, the performance of the bioabsorbable stent is that the bioabsorbable scaffold has a constant high radial strength, minimal recoil, good deliverability, and a controlled rate to the antiluminal cells. It simulates the performance of a substantially durable or non-biodegradable stent in that it has therapeutic drug delivery.

  During phase I, the stent suppresses or prevents the natural pulsation of the blood vessels. While the stent structure suppresses recoil (eg, below 10%) and maintains a circular lumen, the blood vessel reforms and reshapes itself to the diameter of the stent implant, which is positive It corresponds to a re-formation. Early recoil before sufficient formation has occurred may result in negative remodeling. That is, for example, that the stent is molded to a diameter that is significantly less than the original stent implantation diameter of 50% or less of the original deployment diameter.

  At the beginning of Phase II, the radial strength of the stent begins to decline due to the decrease in molecular weight. The radial strength is reduced to a point where the stent can no longer support the vessel wall. As the radial strength of the stent is reduced, the vessel load gradually transitions from the stent to the vessel wall that can ideally support itself with the reshaped diameter. Vascular wall remodeling continues even after the stent loses radial strength. In Phase II, the stent begins to lose mechanical integrity. It is desirable that the structural elements of the stent be incorporated into the vessel wall by the endothelial layer before the stent loses mechanical integrity. The stent is then shattered to allow vasomotion. As the blood vessels move due to vasomotion, the blood vessel wall continues to reform.

  In Phase III, the stent eventually dissolves completely, leaving the treated vessel with an increased diameter that exhibits the same or similar vasomotion as the healthy vessel portion.

  Poly (L-lactide) (PLLA) is attractive as a stent material because of its relatively high strength and rigidity at a human body temperature of about 37 ° C. PLLA has a glass transition temperature of about 60-65 ° C. (Medical Plastics and Biomaterials Magazine, March 1998), so it can maintain stiffness and stiffness at body temperature. This property facilitates the ability of the PLLA stent scaffold to maintain the lumen at or near the deployed diameter without significant recoil (eg, less than 10%).

  In general, the Tg of a semi-crystalline polymer can depend on the morphology and thus the way in which it was processed. Thus, Tg refers to the Tg in which it is associated, such as, for example, PLLA resin, extruded tube, expansion tube, and scaffold Tg.

  The degradation profile generally refers to the time dependence or change of multiple characteristics of a bioabsorbable stent or scaffold over time since implantation in the body lumen of an animal or human patient. This degradation profile may refer to a characteristic change over time in an in vitro test. These characteristics include the molecular weight of the stent body or scaffold polymer, the polymer strength of the stent body or scaffold, the mass of the stent body or scaffold, the mechanical integrity of the stent or scaffold, and the diameter of the stent or scaffold. Directional strength is included.

  Two of the degradation characteristics important for treatment are the time to loss or loss of radial strength, and the time or degradation time of complete stent absorption. The time for loss of radial strength may refer to the time for the stent to maintain radial strength after implantation and to the time after implantation until the stent begins to lose radial strength.

Ideally, when a stent begins to lose radial strength, it is desirable that the bioabsorbable scaffold be absorbed as fast as possible while also meeting all basic safety requirements during its degradation. Such safety requirements include the release of fragments that may cause adverse events such as thrombotic events, or gradual powdering and absorption that does not allow the sudden release of degradation products that cause inflammatory reactions. It is. In this way, the present stent scaffold allows for a positive remodeling of vascular healing and maximizes the benefits described herein of the bioabsorbable scaffold. Therefore, it is very important not only to develop functional and appropriate management methods at the time of embedding (T ₀ ), but also to develop decomposition profile management methods from T ₀ to complete absorption. It is to let you.

  Various embodiments of the present invention include determining characteristics of a bioabsorbable scaffold that provide characteristics of a degradation profile that meets the required or desired degradation characteristics of a specified treatment. The characteristics of the scaffold include an initial value Mn (0) of the number average molecular weight and a degradation rate constant of the scaffold. The inventor of the present application has found that the decomposition rate constant depends on the monomer content of the scaffold, and thus the decomposition rate constant can be managed using the monomer. The characteristics of the degradation profile include the time to radial strength loss of the scaffold and the degradation time (time fully absorbed). Desirable degradation characteristics include the minimum time or opening time of mechanical support and the desired degradation time.

  Preclinical and clinical trials of balloon angioplasty have demonstrated that restenosis is primarily due to early contractile remodeling (vasoconstriction) and is less likely due to hyperplastic healing response . Mintz G, Popma J, Pichard A, Kent K, Satler L, Wong CD, Hong M, Kovach J, Leon M, Circulation 94, 35 (1996); Kimura T, KaTurak. Hamazaki N, Nosaka H, Nobuyoshi M, Mintz G, Popma J, Leon M, Circulation 96, 475 (1997); Di Mario C, Gil R, Camenzin E, Ozine E de Feyter P, Rolandt J, erruys PW, American Journal of Cardiology, 75, 772 (1995); Luo H, Nishioka T, Eigler N, Forrester J, Fishbein M, Berglund H, Siegel R, Arteriosclerosis, Thrombosis and Vascular Biology 16, 1393 (1966). Contractile remodeling can be prevented by embedding a vascular scaffold to keep the vessel open for a period of time. Nobuyoshi et al. Studied the rate of restenosis after angioplasty at 1 month, 3 months, 4 months, 6 months and 1 year. Nobuyoshi M, Kimura T, Nosaka H, Mioka S, Ueno K, Yokoi H, Hamazaki N, Horiuchi H, Ohishi H, Journal of the American C. 12 Using a series of angioplasty, they conclude that the restenosis rate increased significantly and stabilized after 1-3 months after coronary angioplasty. This finding is consistent with the results of Serruys et al. That after balloon angioplasty, restenosis occurs mostly within 3 months and is rarely observed thereafter. Oriston JA, Serruys PW, Circulation: Cadiovascular Interventions 2, 255 (2009); PG, Circulation 77, 361 (1988). Therefore, it is desirable to provide mechanical support to the vessel wall with a bioabsorbable stent for a minimum of 3 months to prevent contractile remodeling and consequent restenosis.

  Thus, for coronary applications, the minimum period for which the stent provides support for positive remodeling (minimum open period) is at least about 3 months. Thus, it is desirable that the time for loss of radial strength or the time for maintaining radial strength be at least about 3 months. For peripheral vascular applications, the minimum patency period should be somewhat longer, eg, at least about 4-5 months. For nasal application, the minimum patency period may be as short as, for example, at least about 3 weeks after frontal sinus surgery with endonasal technique. For nerve applications, the minimum open period may be 3 months.

  With regard to degradation time, the degradation time of a bioabsorbable stent is about 18-26 months for coronary applications, and for peripheral vessels (eg, superficial femoral artery (SFA)) ( Desirably, about 18 months (e.g., 16-20 months), 18-24 months for nerve application, and less than one year for nasal application. Needless to say, the method and its features described herein for managing decomposition profiles are widely applicable and are not limited to the above ranges.

  Various embodiments of the present invention include managing the degradation profile of poly (L-lactide) stents, particularly the time of radial strength loss and degradation time. In these embodiments, the degradation profile is the number average molecular weight at the embedding time or zero time (Mn (0)) and the monomer content (MC) of the finished scaffold, here poly (L-lactide). I.e., MC of L-lactide. Monomer content refers to the content of monomers that are not chemically bonded to the polymer.

  The Mn (0) of the scaffold is the Mn of the polymer scaffold of the final or finished stent product. The final or finished product can refer to a stent or stent scaffold immediately after sterilization, any time after sterilization, or just before or immediately after delivery into the body of a human patient.

  The inventor has discovered through numerous studies that the degradation profile of poly (L-lactide) is governed predominantly by the Mn (0) and degradation rate constant of poly (L-lactide). . As shown below, the inventors have discovered that the degradation rate constant can be controlled by monomer content in a predictable and consistent manner.

It is recognized by the inventor that the desired or required properties of the PLLA scaffold can be predicted using the dynamic degradation properties of PLLA, in particular the dynamic degradation properties of Mn. That is. The inventor has discovered that the degradation profile of Mn of poly (L-lactide) scaffold can be approximated by the following relational equation of autocatalytic kinetics:
In [Mn (t) / Mn (0)] = − kt
Or Mn (t) / Mn (0) = exp (−kt)
Here, k is a decomposition rate constant. C. G. Pitt, M.M. M.M. Gratzl, G.M. L. Kimmel, J .; Surles, A.M. Schindler, Biomaterials 2, 215 (1981).

  The inventor tested the relationship between the lactide content in the PLLA extruded tube and the degradation profile of the PLLA scaffold. FIG. 3A shows the degradation profiles of PLLA scaffolds with different monomer concentrations. FIG. 3B shows the degradation profile corresponding to the above kinetic relevance obtained in an in vitro study of two lots of scaffolds used in animal experiments. All data sets are fully compatible with kinetic relevance.

  With respect to Mn, what has been discovered based on the studies of the present inventors over the past few years is that the Mn of poly (L-lactide) stent scaffolds begins to decrease immediately after the scaffold is implanted. FIG. 4 schematically illustrates the relationship discovered by the inventor that the degradation profile and associated properties (time to radial strength loss and degradation time) depend on Mn and the degradation rate or degradation rate constant. FIG. 4 shows two sets of decomposition profiles corresponding to two initial values Mn (0). Two decomposition profiles are shown for each Mn (0), each with a different decomposition rate or decomposition rate constant. Thus, FIG. 4 shows the effect of Mn (0) and the degradation rate constant on the degradation profile of the bioabsorbable scaffold. For example, if Mn (0) is high, the gradient of the decomposition profile becomes steep due to an increase in the decomposition rate constant, and the decomposition time becomes short. FIG. 4 further shows that the decrease in Mn (0) shifts the decomposition profile in the direction indicated by the arrow, reducing the decomposition time.

  Furthermore, the present inventors have discovered that the time course of radial strength and scaffold integrity during degradation depends on the molecular weight of the scaffold. In general, the values of radial strength and radial stiffness are not a function of the scaffold material alone. Material strength and stiffness can be distinguished from radial strength and radial stiffness because the latter two quantities are stent properties. Stent properties are a complex function determined by the stent material and the stent shape, including the stent pattern and thickness, which are structural elements. Thus, the actual values of radial strength and stiffness depend on the material and stent shape.

  Some of the inventor's studies suggested that the onset of loss of desired mechanical strength (eg, radial strength and tensile strength) can be attributed to the transition molecular weight Mn, Tr of the PLLA backbone. It is related. FIG. 5 is a general graph showing changes in mechanical strength as a function of molecular weight, and is defined by arranging Mn, Tr and Mn, c on the diagram. When the molecular weight exceeds Mn, Tr, the mechanical strength becomes independent of the molecular weight. When the molecular weight is less than Mn, Tr, the mechanical strength begins to decrease, but until the critical molecular weight Mn, c reaches the critical molecular weight Mn, c where the bioabsorbable scaffold becomes brittle and begins to lose mechanical integrity, the mechanical integrity is still Maintained. Since strength reduction is expected to occur prior to loss of mechanical integrity, Mn, Tr is used to minimize the minimum Mn to ensure that the bioabsorbable scaffold maintains adequate strength at the desired onset of degradation. (0) may be predicted.

  In the bioabsorbable PLLA scaffold, Mn and Tr are 47 kDa (Example 4). It is known that Mn and Tr are independent of the decomposition rate constant. The time to reach Mn, Tr coincides with the time of radial strength loss. Mn and Tr are lower limit values of Mn at a desired opening time. If the scaffold Mn falls below Mn, Tr before the desired opening time, the scaffold cannot support the lumen long enough to perform positive remodeling.

  Referring back to FIG. 4, radial strength loss time and decomposition time (Dt) depend on Mn (0) and decomposition rate. When Mn (0) decreases from Mn1 to Mn2, both the radial strength loss time and the decomposition time become shorter. Furthermore, as shown by the profiles of Mn1 and Mn2, the decomposition rate increases, the inclination of the decomposition profile of Mn increases, and the radial strength loss time and decomposition time become shorter.

  The inventor has discovered that when the PLLA scaffold is further degraded to Mn of 30 kDa, the scaffold begins to lose mechanical integrity. Mn at the beginning of the loss of mechanical integrity is denoted as Mn, c.

  As noted above, there is a desired minimum open time for treatment with a stent to provide positive remodeling. Thus, the bioabsorbable scaffold should have a degradation profile with Mn at the desired minimum opening time above Mn, Tr. Mn and Tr represent the lower limit of Mn at a desired minimum opening time. For the treatment of coronary lesions, the minimum open time that meets the basic safety requirements of the scaffold design is about 3 months.

  FIG. 6 shows three degradation profiles of the PLLA scaffold. For example, profile 1 has Mn after 3 months equal to Mn, T, which is acceptable for coronary artery treatment. Profile 2 has the same Mn (0) as profile 1 but has a faster decomposition rate or decomposition rate constant resulting in Mn at the desired opening time shorter than Mn, T. Profile 3 has the same decomposition rate or decomposition rate constant as profiles 1 and 2, but has a smaller Mn (0). As a result, Mn at the desired opening time is smaller than Mn, T. Furthermore, it will be appreciated that a change in the degradation time of the bioabsorbable scaffold will change with either or both of Mn (0) or degradation rate.

  Accordingly, the inventors have discovered that Mn (0) and degradation rate can be adjusted to meet specific therapeutic requirements such as, for example, desired opening time, time of structural integrity loss, and degradation time. This means that a satisfying decomposition profile can be obtained.

  As described above, the present inventors have discovered that the degradation rate constant can be managed in a predictable and consistent manner based on the monomer content of the bioabsorbable scaffold. In particular, the present inventors have discovered that the degradation rate constant showed linearity (linear relationship) with respect to the monomer content of lactide in the PLLA scaffold.

  The inventor has discovered that preclinical studies using porcine models have shown that scaffold integrity tends to depend on kinetic molecular weight gradients in vivo. In addition, the inventor found that in vitro experiments, the onset of radial strength reduction was initially observed in samples associated with higher in vitro degradation rate constants (k). That is. Thus, a clear way to lose molecular weight is essential for managing bioabsorbable scaffold degradation and absorption behavior. The inventor's comparison of in-vivo and in-vitro results showed that the molecular weight data at each stage of the initial degradation was similar between the two models. This discovery is described in the references (Weir N. A., Buchanan F. J., Orr J. F., Diskson G. R. “Degradation of poly-L-lactide. Part 1: in vitro and in vivo discharge.” , Proceedings of the Institute of Mechanical Engineers. Part H: Journal of Engineering in Medicine 218, 307-319 (2004); Hayashi T. in Polymer Science 19, 663-701 (1994)), poly (L-lactide) degradation at an early stage in vivo is mainly caused by simple hydrolysis with minimal enzymatic activity. In line with the discovery that the use of in vitro methods as a surrogate for the degradation behavior of is applicable.

  Lactide is the dominant thermal by-product of the polymer during melt extrusion. By tracking the lactide content at different downstream processing steps of the extruded tube lot, the inventor was convinced that, as shown in Example 1, the extrusion method is the maximum for the lactide content. It was that it contributed to. Thus, the lactide monomer in the resin and the lactide produced during extrusion are primarily or completely the source of monomer in the finished stent scaffold. Furthermore, as shown in Example 2, the present inventors have found that in order to control the lactide content in the finished scaffold, the level of the lactide content of ≦ 0.5 wt% is within the extruded tube. It is sufficient to control the lactide content of the.

The inventor investigated the degradation behavior of lots of extruded tube with different lactide contents by in vitro experiments to test the predictive ability of the kinetic degradation model.

FIG. 3A includes an exponential regression line based on a kinetic model. Each data point represents n = 6 and the error bar represents one standard deviation. R ² (determination coefficient) indicates the model suitability. Decomposition rate constant k is determined according to the model using exponential regression.

Using the autocatalytic model, the decomposition rate constant (k) is calculated for each group in FIG. 3A. FIG. 7 shows the degradation rate constant (k) as a function of lactide content calculated from the linear regression plot of FIG. 3A. FIG. 7 shows the linear direct proportional dependence of the in vitro degradation rate constant (k) on the lactide content. The resulting model (using a sigma plot) is shown in the following relationship:
k (× 10 ⁻³ ) = 10.080 [LA] +1.5131
In the above equation, k is the first-order rate constant (days ⁻¹ ) and [LA] is the lactide content (wt%) in the extruded tube. According to this equation, the higher the initial lactide content in the extruded tube, the faster the scaffold sample will degrade. In addition, linear correlation can be utilized to predict kinetic degradation from a given initial lactide content in the range of about 0.02 wt% to about 1.08 wt%.

  As a result of the various kinetic degradation rates led by different lactide contents, it is expected that the temporal progression of radial strength during degradation will also be affected. By tracking the progress of the radial strength over the entire decomposition time, the present inventor has also shown that the period when the radial strength is maintained in the finished scaffold (FG) is high in lactide content. It was shorter (Example 3).

  FIG. 8 shows the dependence of the degradation profile and its related properties on Mn and monomer concentration. FIG. 8 shows two sets of decomposition profiles corresponding to two initial values Mn (0), namely Mn1 and Mn2. Two degradation profiles are shown for each different Mn (0) corresponding to two different monomer concentrations. Thus, FIG. 8 shows the effect of Mn (0) and monomer concentration on the bioabsorbable scaffold degradation profile. In Mn1 and Mn2, the slope of the decomposition profile becomes steeper as the monomer concentration is higher. FIG. 8 further shows that when Mn (0) is decreased, the decomposition profile shifts downward as indicated by the arrow. Thus, the present inventors have discovered that increasing the L-lactide concentration shortens the period in which the radial strength is maintained by the stent scaffold.

  FIG. 8 can be used to explain how to adjust or select the monomer concentration to obtain the desired radial strength loss time and degradation time. For example, if the required opening time is t1, profile 1 is unacceptable because Mn falls below Mn, Tr and thus loses radial strength before t1. Profiles 2-4 are acceptable because Mn is higher than Mn, Tr at t1, and therefore the time of each radial strength loss occurs after t1. Thus, compared to profile 1, a higher Mn (0) (eg, in profile 3), a lower monomer concentration (eg, in profile 2), or both should be selected or adjusted. Further, the degradation time of the profile 4 can be set longer than that desired for coronary artery treatment, for example, 5 years. In this case, select a low Mn (eg, profile 2), a high monomer concentration (eg, profile 3), or both to obtain a short degradation time while obtaining an acceptable radial strength loss time. Can do.

  The monomer content can be managed in several ways. These methods include the selection of commercially available resins having the desired level of nomer concentration. Furthermore, by controlling the extrusion conditions, monomer production that tends to increase as the extrusion temperature rises can be reduced. Further, for example, in the extrusion step, monomers can be added to the scaffold polymer to increase the monomer concentration.

  While various embodiments of the present invention have been applied to PLLA scaffolds using two different scaffold designs, these methods are generally applicable to other types of bioabsorbable polymers and other scaffold designs. . The method of managing the degradation profile of a bioabsorbable scaffold can be applied to different types of treatments (eg, coronary arteries, SFA, nerves, nose, etc.) and different scaffold designs. Mn (0) and the initial monomer content of the scaffold can be used to manage a degradation profile that meets certain therapeutic specifications. The magnitude of the radial strength supporting the lumen for each case is obtained by the type of polymer selected and the shape of the scaffold (for example, the thickness of the pattern structural element, etc.). As discussed above, there may be a preferred or required degradation profile characteristic of the bioabsorbable stent scaffold that includes the selected degradation time or degradation range, and the time that the scaffold maintains lumen patency. Thus, a method of making a stent scaffold may include determining Mn (0) and / or both that provide the desired degradation profile characteristics. These methods further include making the stent scaffold such that the finished stent scaffold has the determined Mn (0) and MC or both.

Using a predictive model based on the autocatalytic mechanism of PLLA decomposition, the minimum Mn initial value at decomposition time t = 0 can be obtained:
_{lnMn (0) = lnMn, Tr} + k r t (1)
_Where k _r is the standard decomposition rate constant (days ⁻¹ ), Mn (0) is the initial number average molecular weight, and Mn, Tr is the minimum decomposition period t (days) required for product safety. It is the number average molecular weight of mechanical strength transition in (days). The minimum Mn (0) is the minimum initial Mn value of the scaffold that maintains the opening during the desired minimum opening time (eg, 3 months). Each parameter (Mn, Tr, _kr and t) is determined or defined to obtain the predicted Mn (0).

As discussed above, the degradation rate constant and lactide content are represented by the following equation, following linear regression:
k (× 10 ⁻³ ) = 10.080 [LA] +1.5131 (R ² = 0.9988)
In the above equation, k is the decomposition rate constant (days ⁻¹ ), and LA is the initial lactide content (wt%) in the extruded tube. With the regulation of lactide content of ≦ 0.2 wt% in the extruded tube, the fastest possible decomposition rate constant calculated by the above equation is 3.53 × 10 ⁻³ (days ⁻¹ ). It can be seen from Equation 1 that for a given pair of Mn, Tr and t, the faster the decomposition rate constant, the higher Mn (0) is required, so 3.53 × 10 ⁻³ (days ⁻¹ ) is Since it represents the worst case scenario, it is selected as the reference decomposition rate constant (k, r).

Table 1 summarizes the above parameters. Applying these parameters to Equation 1 yields a minimum Mn initial value of 66 kDa. Thus, with a lactide content of 0.02 wt%, the exemplary molecular weight can be Mn (0) ≧ 66 kDa. As explained so far, this molecular weight is considered the sum of the PLLA backbone and the PDLLA coating polymer.

  In other embodiments of determining Mn (0), a method of making a stent scaffold can include determining Mn (0) that provides a desired minimum puncture time. This method involves determining the Mn, Tr of a bioabsorbable stent made from a bioabsorbable polymer, which is about 47 kDa for a PLLA scaffold. The method then includes determining Mn (0) that provides Mn at a desired minimum opening time equal to Mn, Tr. A stent scaffold can be made from a bioabsorbable polymer having a Mn (0) greater than or equal to this determined Mn (0). The determined Mn (0) can be found from a kinetic decomposition model of the bioabsorbable polymer.

  Based on data obtained from long-term in vitro degradation experiments, the Mn of the PLLA scaffold with Mn 110 kDa and the scaffold with an L-lactide monomer content of 0.02% or less in the PLLA backbone are shown in FIG. As shown in the figure, the decomposition time is as long as 29 months.

  Two exemplary modifications include (1) a lactide content of 0.1% or less and (2) a lactide content of 0.2% or less. The selected time at which radial strength is maintained, or Mn (0) that provides time for loss of radial strength, can be determined from the kinetic model.

  In exemplary change (1), Mn (0) of at least 60 kDa should provide radial strength that is maintained for at least 3 months after implantation, and the total degradation time would be expected to be within just 18 months. . In exemplary change (2), Mn (0) of at least 66 kDa should provide the same result. FIG. 10 shows the Mn versus degradation time for two modifications of the PLLA scaffold. (1) 0.1 wt% lactide and Mn (0) = 60 kDa, (2) 0.2 wt% lactide and Mn (0) = 66 kDa.

  In certain embodiments, a method of making a stent can include determining Mn (0) that provides a desired degradation profile characteristic. The determined Mn (0) is the Mn (0) of a polymer having a specific degradation rate or degradation rate constant depending on the monomer content discovered by the present inventors. The Mn (0) therefore corresponds to a given monomer content.

  Depending on these embodiments for determining Mn (0), the desired degradation time or range is selected, and then the stent scaffold Mn (0) or Mn () providing the degradation time or range of the finished stent scaffold. The range of 0) is determined. A stent scaffold can then be made from the bioabsorbable polymer to have a Mn (0) within the determined Mn range.

In these embodiments, the range in which Mn (0) is determined can be determined from a kinetic degradation model of the bioabsorbable polymer. The hydrolysis model for aliphatic polyesters takes the form:
Mn (t) = Mn (0) exp (−kt)
In the above formula, Mn (t) is the number average molecular weight at time t, Mn (0) is the number average molecular weight at t = 0, and k is the hydrolysis rate constant. Pitt, C.I. G. , J. et al. of Applied Polymer Science 26, 3779-3787 (1981); Pitt, C .; G. , Biomaterials 2, 215-220 (1981); Weir, N .; A. , Proceedings of the Institution of Mechanical Engineers, Part H: J. et al. of Engineering in Medicine 218, 307-319 (2004); Weir, N .; A. Part H: J. et al. of Engineering in Medicine 218, 321-330 (2004). The assumptions that are inherent in the model are reasonable on the condition that no mass loss occurs, as mass loss affects the concentration of water and carboxyl end groups in the sample. The above equation can also be rewritten as:
ln [Mn (t) / Mn (0)] = − kt

  Therefore, by expressing the data of Mn (t) / Mn (0) vs. t as a logarithmic-linear plot, the hydrolysis rate constant can be estimated from the slope of the connection point. The degradation rate constant k can be found, for example, from in vitro or in vivo degradation data for a polymer having a given monomer content.

  In certain other embodiments, a method of making a stent can include determining an MC that provides a desired degradation profile characteristic. The determined MC is for a polymer with a specific Mn (0).

  In some of these embodiments for determining the MC, a desired degradation time or range is selected before determining the range of MCs that provides the degradation time range. A stent scaffold is then made from the bioabsorbable polymer so that the MC is within the determined range. The determination range of MC is obtained from a kinetic degradation model of a bioabsorbable polymer. For example, in PLLA, the rate constant k is obtained from Mn (t) / Mn (0) = exp (−kt). MC (0) can then be determined from the in vitro degradation data as shown in FIGS.

  In another embodiment of determining the MC, the desired minimum opening time is selected before determining the Mn, Tr of the bioabsorbable stent. Then, determine MC, Tr that provides Mn at the desired minimum opening time equal to Mn, Tr. A stent scaffold can then be made from the bioabsorbable polymer such that the MC is less than or equal to the determined MC.

  The determined MC can be obtained using a kinetic degradation model of the bioabsorbable polymer. For example, in PLLA, the rate constant k is obtained from Mn (t) / Mn (0) = exp (−kt). The MC can then be determined from the in vitro degradation data as shown in FIGS.

  In the embodiment described above, Mn (0) or MC can be determined for a bioabsorbable polymer scaffold that provides degradation profile parameters, and then a stent scaffold with Mn (0) and MC can be made . Embodiments of the invention include making a stent scaffold with determined Mn (0) and MC.

  In extrusion, the polymer is processed above its melting point (Tm). Since the viscosity of the molten polymer increases with temperature, the higher the Mn of the resin, the higher the temperature required for processing in the extruder. However, monomer production increases with temperature, and the decrease in Mn increases with temperature. An exemplary melt treatment of PLLA resin can be performed on a 3/4 ″ single screw extruder. For resins with Mn of about 350 kDa, the treatment temperature is 200-210 ° C. and the residence time is 8-10 minutes. As the tube exits the die, it is quenched in a room temperature water bath, the barrel pressure of the extruder is about 2000 psi, and the crystallinity after extrusion is about 10-15%.

  When applied to coronary arteries, the outer diameter of the polymer tube used in stent manufacture can be 2-4 mm. The outer diameter when applied to SFA is even larger, for example, 4 to 9 mm. The diameter may exceed these ranges. Although the wall thickness of the polymer tube can be 0.05 to 3 mm, the present invention can be applied to a tube having a wall thickness of less than 0.05 mm and a tube having a wall thickness of more than 3 mm.

  Prior to laser cutting, the radial strength of the stent may be increased by expanding the tube in the radial direction and increasing its radial strength. The tube can also be stretched or stretched axially during the expansion process. The radial expansion process tends to favorably align the polymer chains along the hoop direction, resulting in enhanced radial strength. The radial expansion step is essential to create a stent scaffold with strong and thin struts sufficient to support the lumen when implanted.

  The tube is radially expanded by being heated to a temperature between the Tg and melting point of the polymer. When the tube is expanded, it is cooled below the Tg of the polymer, generally to room temperature, to maintain the tube at its expanded diameter. The tube is expanded and then cooled at a non-equilibrium rate, thereby maintaining the tube at the expanded diameter. The rate of radial expansion can be 200-500%. The ratio of radial expansion is defined as RE% = (RE ratio−1) × 100%, where RE ratio = (expansion tube inner diameter) / (original tube inner diameter). The percentage of axial expansion experienced by the polymer tube is defined as AE% = (AE ratio-1) × 100%, where AE ratio = (expansion tube length) / (original tube length). .

The tube can be expanded in the radial direction by blow molding a tube placed inside a glass mold. The tube is heated and expanded to the inner diameter of the mold. For example, if warm air is blown from the heating nozzle into the mold while the heating nozzle is translated along the length of the mold, the tube is expanded along with the parallel movement of the nozzle. Here, axial extension may be applied in a state where axial tension is applied to the tube. The exemplary embodiment tube is expanded from an inner diameter of 0.018 ″ / outer diameter of 0,056 ″ to an inner diameter of 0.072 ″ / outer diameter of 0.084 ″, with a radial expansion (RE) of 350%, longitudinal Stretching is 50%. Here, RE = [(outer diameter) _finish / (outer diameter) _start −1] × 100. An exemplary PLLA tube may be heated at about 70-110 ° C. during its expansion.

  A stent pattern is cut on the expanded tube by, for example, laser processing. Tube expansion reduces the wall thickness of the tube. In the case of a coronary stent, the width and thickness can be, for example, 140 to 160 μm. In the case of SFA, the width and thickness may be 180 to 230 μm.

  After cutting the stent pattern into the expanded tube, the stent scaffold may optionally be coated with a drug delivery coating that may include polymers and drugs. Exemplary stents may include a PLLA scaffold and a coating composed of, for example, a 1: 1 (by weight) poly (DL-lactide) and everolimus.

  In order to be able to deliver the stent at any time, the stent is secured to the delivery balloon. This process compresses or crimps the stent onto the balloon with a reduced diameter. In an exemplary embodiment, the stent is a multi-step process with a dwell (stop) period between each diameter reduction, from the cutting diameter to the crimping diameter (e.g., 0.136 "to 0.047). ") To be crimped. The crimping temperature of the stent can be above ambient temperature, for example about 48 ° C. or slightly below the TLA of PLLA. A sheath may be placed on the stent immediately after crimping to prevent recoil. The stent may then be placed in a sealed bag.

  Then, after crimping and packaging the crimped stent, the stent may be subjected to a final sterilization treatment. Terminal sterilization refers to the final sterilization step in stent manufacture, which exposes the stent to radiation such as, for example, electron beams or gamma rays. Typically, for example, the stent is sterilized in one step, with one or more passes through the radiation without any other steps. Therefore, the final irradiation step may be only the sterilization step. No radiation exposure is added after final sterilization. Terminal sterilization is usually performed on the stent after crimping and packaging, but may be performed before one or both of crimping and packaging.

By sterilizing the packaged stent and catheter, the bioburden of the stent and delivery system is reduced to a prescribed level. Bioburden generally refers to the number of microorganisms that contaminate a subject. The degree of sterilization is typically measured by a sterility assurance level (SAL), which refers to the probability of viable microorganisms present in the product unit after sterilization. The SAL required for a product depends on the intended use of the product. For example, products such as stents used in the body fluid pathway are considered Class III devices and require a SAL of ^10-6 . SALs for various medical devices are found in materials from the American Association for Medical Device Promotion (AAMI) in Arlington, VA (VA).

  Sterilization can be performed by exposing the stent and catheter to radiation such as electron beam (e-beam), gamma radiation, and X-ray sterilization. The radiation dose for sterilization can be determined by selecting the radiation dose that provides the required SAL. The sample can be exposed to the required dose by passing it one or more times. An exemplary radiation dose for stent sterilization may be 20-35 kGy.

  The resin has a molecular weight Mn, r and a monomer content MC, r before entering any processing step. As mentioned earlier, both Mn and MC change during the manufacturing process. Mn is significantly reduced during extrusion and radiation sterilization. The higher the extrusion temperature, the more Mn decreases. As the radiation dose increases, Mn decreases greatly. For example, a PLLA resin having an Mn of 350 kDa has an Mn of an extruded tube of 250 kDa due to an extrusion temperature of 215 ° C. Prior to electron beam sterilization, the MLA = 250 kDa PLLA stent scaffold is reduced to 90-100 kDa Mn after exposure to a radiation dose of 27.5 kDa.

  As noted above, MC can increase during extrusion. The higher the extrusion temperature, the more monomer production. Mn, r, MC, r in combination with processing parameters, particularly extrusion and radiation sterilization processing parameters, may not provide Mn (0) with the desired degradation profile.

  For example, Mn, r, MC, r and process parameters may result in a stent scaffold having Mn (0) above Mn (0) that provides the desired degradation profile. That is, the time for maintaining the radial strength and / or the decomposition time is longer than the desired time. For example, a scaffold made from PLLA resin with Mn = 365 kDa and LLA having a monomer content of about 0.1% is processed using the exemplary processing conditions disclosed above to provide Mn = 100-110 kDa. Create the final PLLA scaffold. The degradation time of this stent scaffold is about 2.5-3 years and is not unacceptable, but a shorter time would be desirable for coronary artery and other applications.

  If Mn (0) is greater than the desired value, certain embodiments can be adjusted during processing to provide Mn (0) that provides the desired processing parameters. In particular, the Mn of the bioabsorbable polymer can be reduced by exposure to radiation, such as an electron beam or gamma radiation. In the method of changing the molecular weight by using the electron beam irradiation, the molecular weight is changed by changing the electron beam irradiation amount to the PLLA scaffold, and the desired molecular weight or Mn of the target scaffold is obtained. The effect of the electron beam depends on PLLA molecular chain scission that reduces the molecular weight. When the initial value of the molecular weight of the scaffold is found, there is a relationship between the electron beam irradiation amount and the obtained molecular weight within a range in which the molecular weight can be managed using the electron beam. Thus, a desired broad range of starting molecular weight (below the initial molecular weight) can be obtained by varying the electron beam dose on the scaffold.

  In the above embodiments, several methods are disclosed for determining the desired degradation profile, in particular Mn (0), which provides the scaffold degradation and opening times. With respect to the opening time, Mn (0) is determined which provides the time for radial strength loss in the desired opening time range.

  In certain embodiments, electron beam irradiation is used to adjust Mn and to find the desired decomposition characteristics, such as the time to maintain the radial strength while the scaffold is decomposed, and thus the decomposition time. Mn (0) obtained can be obtained.

  In some embodiments, the radiation dose in the final radiation sterilization step is adjusted to reduce the crimped Mn to Mn (0) that provides the desired degradation profile. In this embodiment, the radiation dose required to reduce the Mn after crimping to the desired Mn (0) may be determined from the relationship between Mn and the radiation dose determined experimentally or empirically. . In an exemplary embodiment, the initial Mn value of the PLLA scaffold after crimping is about 250 kDa. The time of loss of radial strength may be shortened to 2.5-3 months by determining Mn (0) corresponding to this time of loss, as disclosed above. The radiation dose required to reduce the Mn after crimping to this Mn (0) is then obtained from the relationship of radiation dose to Mn. For example, the desired Mn (0) can be achieved by adjusting the electron beam radiation dose from 31 kGy to 75 kGy.

  The inventor has determined that a scaffold with Mn (0) of about 70 kDa provides a degradation time of less than 18 months and the time for radial strength loss is at least 3 months. In this way, the radiation dose in the final sterilization process can be adjusted to reduce Mn to about 70 kDa.

  In other embodiments, in addition to the final radiation step, the Mn is adjusted by radiation at one or more points in the manufacturing process to obtain a scaffold with Mn (0) that provides the desired degradation profile. Can do. The advantage of adjusting the molecular weight in two or more steps is that each exposure can be less than a single exposure compared to a single exposure in the final sterilization process. Low radiation doses will reduce the possibility of adverse effects from high radiation doses, such as high temperatures at which the polymer of the scaffold changes the mechanical properties of the scaffold.

  In some embodiments, the stent scaffold is exposed to radiation only once at other points during manufacture. In other embodiments, the stent scaffold is exposed to radiation at two or more points during manufacture in addition to the final sterilization process. In these embodiments, the stent scaffold is pre-crimped after laser cutting, applied and post-crimped after laser cutting, post-crimped before crimping, expanded and before laser cutting, or extruded and before expanded. Can be exposed to radiation.

  In some of these embodiments, the radiation dose in the final sterilization process is 20-31 kGy. Therefore, the radiation dose between the manufacturing steps before the final sterilization treatment is adjusted so that the Mn of the scaffold after the final sterilization treatment becomes a desired Mn (0). In an exemplary embodiment, in addition to the 20-35 kGy terminal sterilization process, an additional radiation exposure step is performed only between laser cutting and crimping. The desired Mn (0) after final sterilization may be obtained by adjusting the radiation dose between laser cutting and crimping. This radiation dose may be 6 to 50 kGy. The radiation dose mainly depends on Mn and monomer content. For example, in the case of a PLLA scaffold with a Mn of about 250 kDa and a monomer content of about 0.1 wt% after laser cutting, the radiation dose is adjusted to lose the degradation time of about 18 months and the radial strength in about 3 months The scaffold after the final sterilization treatment of 60 kDa can be obtained.

  In certain embodiments, Mn may be adjusted (decreased) in a pre-decomposition process of the PLLA scaffold to obtain Mn (0) that provides the desired decomposition characteristics. The method includes adjusting the PLLA scaffold degradation characteristics by pre-degrading the scaffold with hydrolysis. The pre-decomposition process reduces the Mn of the scaffold. The method may include exposing the scaffold to a fluid for a period of time causing the scaffold polymer to hydrolyze. In such embodiments, the scaffold may be exposed to the pre-decomposition processing fluid by spraying the scaffold with fluid or immersing the scaffold in the fluid.

  The pre-cracking treatment fluid may be water or an aqueous fluid. The predegradation treatment fluid may be a simulated fluid that is as close as possible to the in vivo degradation of the scaffold. A fluid that causes only a slight change in pH, such as a buffer solution, may be used when the scaffold is pre-degraded. The buffer solution is an aqueous solution containing a mixture of a weak acid and its conjugate base, or a mixture of a weak base and its conjugate acid. Buffer solutions have the property that the pH of the solution hardly changes even when a small amount of strong acid or base is added to the solution. This is important because hydrolytically degradable aliphatic polymers, such as PLLA, have acidic degradation products that can cause a decrease in the pH of the polymer under degradation or the local pH of the polymer. Buffer solutions are used as a means of keeping the pH almost constant in a variety of chemical applications.

  In some embodiments, a phosphate buffered saline (PBS) solution is used. PBS is an aqueous saline solution containing sodium chloride, sodium phosphate, and (depending on formulation) potassium chloride and potassium phosphate.

  Further, in such embodiments, the scaffold may be pre-decomposed after selected steps of the manufacturing process. In some embodiments, the scaffold is pre-decomposed after laser cutting and before application. In other embodiments, the scaffold is pre-decomposed after application and before crimping.

  In other embodiments, a pre-disassembly process can be performed prior to laser cutting. In this embodiment, if the commercially available resin does not meet the molecular weight design requirements, a pre-decomposition treatment of the extruded tube, expanded tube, or PLLA resin prior to extrusion may be performed.

  In some embodiments, the pre-decomposition treatment with the fluid can be performed at any temperature between room temperature, that is, between 20-30 ° C. (including both-side temperatures). In other embodiments, the pre-decomposition process may be performed at a temperature below room temperature or above room temperature. The pre-decomposition process at the selected temperature is performed by exposing the scaffold, tube or resin to a fluid at the selected temperature. Since the decomposition increases as the temperature rises, the time for the pre-decomposition treatment can be shortened if it is carried out at a temperature higher than room temperature. Depending on the embodiment, the temperature of the pre-decomposition treatment can be 40-70 ° C or 50-70 ° C. More narrowly, the pre-decomposition treatment temperature is 40-45 ° C, 40-70 ° C, 45-50 ° C, 50-55 ° C, 55-60 ° C, 60-65 ° C, or 65-70 ° C.

  After the pre-decomposition process, the scaffold, tube or resin is removed from the pre-decomposition process fluid and washed with water to remove any remaining pre-decomposition process fluid or salt. The washed scaffold may then be dried in vacuum.

  The pre-degraded scaffold, tube or resin is pre-degraded to the target Mn providing Mn (0) that provides the desired degradation time and / or time for radial strength loss. Therefore, the determination of the target Mn of the pre-decomposed resin must take into account the reduction in Mn throughout the manufacturing process, especially throughout extrusion and sterilization. Determining the target Mn of the pre-decomposed extruded or expanded tube must take into account the reduction in Mn throughout other manufacturing processes, especially throughout sterilization. Determining the target Mn of the scaffold to be pre-decomposed must take into account the reduction in Mn throughout other manufacturing processes, especially sterilization. Since the decomposition rate depends on the temperature, the time of the predecomposition process depends on the temperature of the predecomposition process.

  The inventor has found through experiments that the scaffold degradation mechanism remains unchanged even in the higher temperature range compared to the degradation mechanism at about 37 ° C. in the human body. FIG. 11 shows logk (degradation rate constant) vs. 1 / T within the relevant temperature range for PLLA samples that are degraded in PBS buffer solution. The good data fit to the straight line shows that the kinetic decomposition model is fit at all test temperatures. Thus, once deployed in a patient, a pre-degraded PLLA scaffold product is expected to be degraded in the same manner as a non-pre-degraded scaffold.

  FIG. 12 shows Mn (t) versus pre-decomposition processing time at five temperatures. M (t) is normalized by the molecular weight before the predecomposition process. The pre-decomposition processing time required to achieve the desired initial molecular weight can be evaluated from FIG. For example, to reduce the molecular weight (Mn) of a pre-sterilized scaffold from 250 kDa to 100 kDa, the laser-treated scaffold can be pre-decomposed in PBS buffer solution at _70 ° C. for about 5 days. It will be good.

  The embodiments described above include adjusting the molecular weight of the finished scaffold with radiation to obtain the desired degradation characteristics. As already explained, the degradation time of a PLLA scaffold made from a resin of about 350 kDa and a monomer concentration of 0.1 wt% and treated as described above is 2.5-3 years, such as coronary artery treatment A Mn (0) of about 100-110 kDa higher than the Mn (0) desired for the application is obtained.

  The scaffold further includes a drug delivery coating on the scaffold that includes a bioabsorbable polymer and a drug such as an antiproliferative agent. The bioabsorbable polymer acts as a drug carrier and controls drug release. The purpose of this drug is to reduce smooth cell proliferation that occurs during the period of 1-3 months after implantation of the scaffold. The drug release profile (drug release volume versus time, or cumulative drug release versus time, duration of drug release) is managed, adjusted or controlled by absorption of the polymer carrier or coating.

  For any bioabsorbable polymer, the degradation profile of the polymer carrier depends on the initial molecular weight of the polymer. The lower the starting molecular weight of the polymer carrier, the shorter the release time of the antiproliferative agent. In general, the release profile of any drug in the carrier depends on the starting molecular weight of the carrier. In general, there is a minimum range of drug release times for treating smooth cell proliferation, about 1-3 months. Therefore, the drug release profile of the coating should meet this limitation.

  After coating the scaffold, any radiation exposure to the stent scaffold not only reduces the molecular weight of the scaffold but also reduces the molecular weight of the polymer carrier of the coating. The decrease in the molecular weight of the carrier affects the drug release profile. In particular, since the decomposition time of the carrier is shortened, the drug release period is shortened. Accordingly, the molecular weight (Mn) of the polymer carrier prior to any radiation exposure should be selected such that after any radiation exposure, Mn is high enough to provide the desired minimum drug release time.

  In other embodiments, both the scaffold PLLA resin and the coated PDLLA resin may be selected to provide a desired degradation profile. In particular, these resins are selected such that the scaffold and the coating Mn (0) each provide the desired degradation profile. In addition, these resins are selected so that the monomer content of the scaffold and coating also provide the desired degradation profile, respectively.

  Returning to the previous example, a PLLA resin with a high Mn of about 350 kDa and a monomer content of up to 0.1% for scaffold production is 100-110 Mn with a degradation time of 2.5-3 years. To produce a scaffold having after sterilization. A coating comprising a polymer carrier can be made from a PDLLA resin having a Mn of about 47 kDa and a monomer content of up to 4%. Polymer carrier mass loss may begin to occur within only a few weeks after implantation.

  It may be desirable to shorten the degradation time of the scaffold and increase the degradation time of the coating. In some embodiments, the Mn and monomer contents of both the PLLA resin for the scaffold and the PDLLA resin for the coating are adjusted or selected to obtain the desired degradation profile, such as degradation time, of the scaffold and coating.

  As discussed above, Mn (0) and MC can be determined to correspond to the desired decomposition profile. Accordingly, Mn (0) and MC corresponding to the desired decomposition profile can be obtained by changing the Mn and monomer contents of the resin. The resulting Mn (0) and MC correspond to specific processing conditions because Mn is affected by extrusion conditions and radiation dose and monomer content is affected by extrusion conditions. In a further embodiment, the extrusion conditions, primarily temperature, can be adjusted to obtain the desired degradation profile.

  For example, the desired degradation profile of the scaffold includes a degradation time of 16-20 months. The desired degradation time of the PDLLA coating polymer is 2-3 months. These degradation characteristics are obtained with PLLA scaffolds made from resins with Mn (0) of 150-200 kDa and PDLLA coatings with Mn (0) of 70-100 kDa.

[Stent structure / characteristic data for in vitro experiments]
The stent scaffold pattern used in the in vitro experiments described herein corresponds to the pattern described in Yang & Jow et al. US Patent Application No. 12 / 447,758 (US2010 / 0004735). The cross-section of the scaffold strut is 150 × 150μ.

Example 1 Demonstration that Extrusion Molding Contributed Maximum to Lactide Content
FIG. 13 shows the effect of the PLLA scaffold manufacturing process on monomer lactide production. Two lots of extruded tubes were made with trace concentrations (<0.02 wt%) and high concentrations (0.97 ± 0.03 wt%). At trace concentrations of lactide (<0.02 wt%), a slight increase in lactide content was detected from the extruded tube to the finished product (FG). This is due to the lactide content in the PDLLA coating polymer and is not expected to contribute to PLLA scaffold degradation. The reason is that lactide should be eluted when it comes into contact with water, considering that the PDLLA coating is thin and the solubility of lactide in water is high. Therefore, the lactide content in the extruded tube represents the lactide content in the finished product (FG). For a lot of extruded tubes with a high lactide content (0.97 ± 0.03 wt%), a slight decrease was observed from the extruded tube to the finished product FG. This decrease can be attributed to an increase in the probability of ester bond cleavage in the cyclic lactide monomer due to the energy of the electron beam. This should produce other types of low molecular weight species, such as di-lactic acid, which have equivalent effects on degradation. Such a phenomenon is not observed at low levels, such as the proposed limit of lactide content. In this case, the extruded tube represents the worst case lactide content compared to the corresponding finished product FG.

[Example 2 shows that the lactide content in the extruded tube was equivalent to the content in the finished product FG]
Various lactide content levels (0.02 (trace), 0.17, 0.57, and 1.08 wt) obtained by spiking (adding) a predetermined amount of L-lactide to the PLLA resin before extrusion Four groups of PLLA scaffold finished products (FG) were made from lots of extruded tube (% lactide). N = 10 for all finished product groups, n = 2 for groups “1.08 wt%” and “0.02 wt%” and n for group “0.17 wt%” for extruded tube groups = 10, n = 21 for the group “0.57 wt%”. Error bars represent one standard deviation.

  FIG. 14 shows the lactide content of an extruded tube obtained from gas chromatography by a flame ionization type detection method. FIG. 14 shows that either the lactide content in the extruded tube is equivalent to the lactide content in the FG or represents a worst case scenario in terms of degradation impact. That is. Further, FIG. 14 shows that the loss of lactide content from the extruded tube to the FG decreases with decreasing lactide content in the extruded tube. If the extruded tube contains less than about 0.5 wt% lactide, either the lactide content between the extruded tube and FG remains either unchanged or slightly increases in FG, but again This is due to the known amount of lactide in the PDLLA coating polymer. Therefore, by managing the lactide content in the extruded tube at the 0.5 wt% lactide content level, the lactide content in the FG can be sufficiently managed. This data shows that loss of lactide content at high levels generates other types of low molecular weight species that are predicted to increase the probability of lactide molecules being cleaved, i.e., have the same effect on degradation. , Shows what is happening. Accordingly, the lactide content of the extruded tube is determined to be equivalent to the corresponding lactide content in the FG in view of the influence on decomposition.

[Example 3—Lactide mixing in extrusion: Influence on radial strength change during decomposition in vitro]
FIG. 15 shows the course of radial strength throughout the degradation as a function of the lactide content of the four lots of extruded tubing from Example 2. Each data point represents n = 6. Error bars represent one standard deviation. The course of radial strength over the entire decomposition time was followed. FIG. 15 shows that the higher the lactide content, the shorter the period in which the radial strength is maintained in the FG. Although this effect was not demonstrated at lactide content levels of “0.02 wt%” and “0.17 wt%” in this experiment for a limited period of time, similar results indicate that these lactide content levels But it is expected to be observed.

  Although this effect was not demonstrated at lactide content levels of “0.02 wt%” and “0.17 wt%” in this experiment for a limited period of time, similar results indicate that these lactide content levels But it is expected to be observed.

[Example 4-Determination of Mn and Tr of PLLA scaffold]
Table 2 is a summary of two experiments used to determine the Mn, Tr of the PLLA scaffold. As shown in Table 2, each experiment concludes the range of Mn and _Tr .

  A radial strength test of FG with a lactide content of 0.95 wt% resulted in Mn, Tr falling in the range of 52 kDa to 45 kDa. This is in line with the upper limit of 51 kDa obtained from a tensile test on the dogbone in the circumferential direction. More details on Mn, Tr by demonstrating that in vitro experiments on FG with 0.51 wt% lactide content, the degraded scaffold with Mn of 47 kDa can still maintain high radial strength Analyzed. Therefore, 47 kDa is selected as Mn, Tr of the PLLA bioabsorbable scaffold.

  While particular embodiments of the present invention have been shown and described, it will be appreciated by those skilled in the art that changes and modifications can be made in a broader manner without departing from the invention. Accordingly, all such changes and modifications should be included within the scope of the claims to fall within the true spirit and scope of the present invention.

Claims (32)

Providing a bioabsorbable polymer;
Selecting a desired degradation time range of the bioabsorbable implantable stent scaffold so that it is completely absorbed after implantation;
Determining the range of Mn (0) of the stent made from the bioabsorbable polymer, wherein the completed stent provides the range of degradation time, wherein the determined Mn (0) of the completed stent is ) Is determined from a kinetic degradation model of the bioabsorbable polymer;
Creating a stent scaffold from the bioabsorbable polymer, the stent scaffold having a Mn (0) within the determined range of Mn (0);
A method of manufacturing a bioabsorbable stent. Adjusting the molecular weight of the bioabsorbable polymer during the step of making the stent scaffold such that the stent scaffold has the Mn (0) within the determined range of the Mn (0). Including further;
A method of manufacturing the stent of claim 1. The molecular weight is adjusted by pre-degradation by exposure to radiation or hydrolysis of the bioabsorbable polymer,
A method of manufacturing the stent of claim 2. The bioabsorbable polymer is PLLA;
A method of manufacturing the stent of claim 1. The bioabsorbable polymer is PLLA;
The kinetic model has the form Mn (t) / Mn (0) = exp (−kt), where t is the decomposition time, Mn (t) is Mn as a function of time, and k is the decomposition rate. A constant,
A method of manufacturing the stent of claim 1. The range of the decomposition time is 13 to 16 months, 16 to 20 months, or 20 to 24 months.
A method of manufacturing the stent of claim 1. Providing a bioabsorbable polymer;
Selecting a desired minimum open time of the bioabsorbable implantable stent scaffold to provide to the implant site;
Determining Mn at the time of loss of radial strength of a bioabsorbable stent scaffold made from the bioabsorbable polymer;
Determining Mn (0) of a stent scaffold made from the bioabsorbable polymer that provides Mn at a desired minimum open time equal to Mn at the time of loss of radial strength, Said determining step wherein said Mn (0) is determined from a kinetic degradation model of said bioabsorbable polymer;
Creating the stent scaffold from the bioabsorbable polymer, wherein the stent scaffold has a Mn (0) greater than or equal to the determined Mn (0);
A method of manufacturing a bioabsorbable stent. Adjusting the molecular weight of the bioabsorbable polymer during the step of making the stent scaffold such that the stent scaffold has a Mn (0) that is greater than or equal to the determined Mn (0). ,
A method for manufacturing the stent of claim 7. The molecular weight is adjusted by pre-degradation by exposure to radiation or hydrolysis of the bioabsorbable polymer,
A method of manufacturing the stent of claim 8. The bioabsorbable polymer is PLLA;
The Mn in the radial strength loss is about 47 kDa.
A method for manufacturing the stent of claim 7. The bioabsorbable polymer is PLLA;
The kinetic model has the form Mn (t) / Mn (0) = exp (−kt), where t is the decomposition time, Mn (t) is Mn as a function of time, and k is the decomposition rate. A constant,
A method for manufacturing the stent of claim 7. The desired minimum opening time is 3 months;
A method for manufacturing the stent of claim 7. Providing a bioabsorbable polymer, wherein the bioabsorbable polymer is formed by a polymerization reaction of monomers and is made of repeating units;
Selecting a desired degradation time range of the bioabsorbable implantable stent scaffold to be completely absorbed after implantation;
Determining a range of monomer content within the bioabsorbable polymer to provide a range of degradation times for the stent scaffold, the determined range of monomer content being determined by the bioabsorption Said determining step determined from a kinetic degradation model of the functional polymer;
Making the stent scaffold from the bioabsorbable polymer, the stent scaffold having the monomer content within a determined range;
A method of manufacturing a bioabsorbable stent. Adjusting the monomer content during the creating step so that the stent scaffold to be produced has the monomer content within a determined range;
A method for manufacturing the stent of claim 13. The monomer content is adjusted by monomer addition before the extrusion step, temperature control of the extrusion step, or a combination thereof.
15. A method of manufacturing the stent of claim 14. The bioabsorbable polymer is PLLA;
A method for manufacturing the stent of claim 13. The range of the decomposition time is 13 to 16 months, 16 to 20 months, or 20 to 24 months.
A method for manufacturing the stent of claim 13. Providing a bioabsorbable polymer, wherein the bioabsorbable polymer is made of repeating units formed by a monomer polymerization reaction;
Selecting a desired minimum open time of the bioabsorbable implantable stent scaffold;
Determining Mn at the time of loss of radial strength of the bioabsorbable stent scaffold made from the bioabsorbable polymer;
Determining the monomer content of the bioabsorbable polymer of a finished stent that provides Mn at a desired minimum opening time equal to Mn at the time of loss of radial strength, wherein the determined monomer content is Said determining step determined from a model of degradation characteristics of said bioabsorbable polymer versus monomer content;
Creating the stent scaffold from the bioabsorbable polymer, wherein the bioabsorbable polymer of the stent scaffold has a monomer content less than or equal to the determined monomer content. And comprising:
A method of manufacturing a bioabsorbable stent. Providing a bioabsorbable polymer scaffold made of PLLA prior to the radiation exposure step, wherein the MLA of the PLLA of the scaffold is at least about 250 kDa;
Selecting a desired minimum open time of the PLLA scaffold to provide to an implantation site;
Providing Mn upon loss of radial strength during decomposition of the PLLA scaffold;
Determining Mn (0) of the PLLA scaffold that provides the Mn of the PLLA scaffold at the desired minimum opening time equal to Mn at the loss of radial strength;
Performing sterilization, including exposing the PLLA scaffold to a radiation dose of 31-75 kGy that reduces the Mn of the PLLA scaffold to a value that is not less than Mn (0);
A method of manufacturing a bioabsorbable stent. The determined Mn (0) is determined from a kinetic degradation model of the bioabsorbable polymer.
The method of claim 19. Mn in the radial strength loss is about 47 kDa,
The method of claim 19. The time to loss of radial strength of the provided scaffold is reduced from about 12 months to 3 months,
The method of claim 19. The radiation dose reduces Mn of the PLLA scaffold to the Mn (0),
The method of claim 19. The monomer content of the sterilized scaffold is 0.2 wt% or less,
The radiation dose reduces Mn of the PLLA scaffold to 66 kDa,
24. The method of claim 23. Providing a PLLA polymer scaffold, wherein the PLLA polymer tube has a Mn of at least 250 kDa;
Exposing the laser-cut scaffold to a first radiation dose before crimping to reduce Mn;
Crimping the exposed scaffold to a reduced diameter on a delivery balloon;
Exposing the crimped scaffold to a second radiation dose of 20-31 kGy for sterilization to reduce the Mn to Mn (0), wherein the Mn (0) is 16-20 months Said exposing step providing a degradation time and a time of radial strength loss of at least about 3 months;
A method of manufacturing a bioabsorbable stent. The monomer content of the sterilized scaffold is 0-0.1 wt%, 0.1-0.2 wt%, 0.2-0.5 wt%, 0.5-1 wt%, and after final sterilization The Mn is 55 to 110 kDa,
The method of claim 19. The first radiation dose is 6 to 50 kGy.
The method of claim 19. Providing a bioabsorbable polymer scaffold made of PLLA prior to the radiation exposure step, wherein the MLA of the PLLA of the scaffold is at least about 250 kDa;
Exposing the scaffold to radiation for sterilization, wherein the radiation reduces the Mn of the scaffold to 70 kDa or less, and the Mn of the exposed scaffold is less than 18 months of the exposed scaffold. Fold degradation time and said exposing step providing time for at least 3 months to radial strength loss;
A method of manufacturing a bioabsorbable stent. The monomer concentration of the scaffold is 0.2 wt% or less;
30. The method of claim 28. Providing a PLLA resin having a Mn of 150-200 kDa;
Processing the PLLA to form a PLLA scaffold;
Forming a coating comprising PDLLA having a Mn of 80-100 kDa on the PLLA scaffold;
Exposing the coated scaffold to radiation for sterilization, the exposing step reducing Mn of the PLLA scaffold to 70 kDa or less by radiation exposure;
A method of manufacturing a stent. Providing a bioabsorbable polymer resin;
Extruding the polymer resin to form a tube;
Radially expanding the polymer tube;
Creating a stent scaffold from the expanded tube;
Radiation sterilizing the scaffold;
Pre-degrading at least one of the resin, the extruded tube and the radially expanded tube by hydrolysis to reduce Mn;
A method of manufacturing a stent. Creating a PLLA stent scaffold made of a biodegradable stent scaffold, wherein the MLA of the PLLA stent scaffold exceeds 250 kDa;
Pre-degrading the PLLA stent scaffold with hydrolysis prior to radiation sterilization to reduce the Mn of the scaffold to 100 kDa or less, wherein the pre-degrading process comprises 18 months Providing a degradation time of the scaffold of less than the degradation process;
A method of manufacturing a stent.

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